JP2005521440A - Bioresorbable osteoconductive composition for bone regeneration - Google Patents

Abstract

Disclosed are bioresorbable bone conduction compositions and methods of using the compositions as scaffolds for bone repair in periodontal, alveolar or maxillary regeneration, skull defects, and spinal cord regeneration. The bioresorbable composition includes a bioresorbable polymer, a microparticulate or nanoparticulate filler, and a pore generating material. The bioresorbable polymer can be electrically unsaturated and crosslinkable with a crosslinker. The micro- or nano-filler can be any natural biocompatible material such as metal, calcium carbonate, carbon, biocompatible synthetic material, or bioceramic such as hydroxyapatite. The pore-generating material can be a blowing agent such as carbonate and acid.

Description

(Background of the Invention)
The Government of the United States, Joseph D. We have the rights in this invention by NIH / NIDR grant number 1 R43 DE 12290-01A1 to Gresser and NIH / NIAMS grant number AR 45062 to Kai-Uwe Lewandrowski.

  U.S. filed on Feb. 5, 2002. S. S. N. Claim priority to 60 / 354,833.

  The present application generally relates to the use of bioresorbable osteoconductive compositions for bone repair and scaffolds formed therefrom. In a preferred embodiment, the present application provides a living organism comprising a micro- or nano-filler and a pore-forming agent for mouth reconstruction, skull defect and spinal cord repair such as periodontal, alveolar or maxillary bone regeneration. It relates to the use of the resorbable osteoconductive composition.

  Studies in the mid 1980s estimated that approximately 4.5 million people in the United States suffer from fractures each year. Bone diseases such as periodontal disease that cause bone loss are added to this problem. Accordingly, bone repair materials are actively sought for bone repair and regeneration. Biodegradable and biocompatible polymer compositions are useful for bone grafting purposes, bone repair purposes, bone replacement purposes, or bone implant fixation purposes.

  Many problems exist with current bone grafting methods, bone repair methods, and bone implant fixation methods. In periodontal reconstruction or implant fixation, the bone void needs to be filled with graft material that supports the structural integrity of this site throughout the course of new bone regeneration. The currently available modality for the treatment of periodontal disease with subsequent bone loss improves its progression but has only minimal potential to regenerate the dental support device.

  This is also true in the reconstruction of the maxilla and mandible. Autografts and allografts are used in current bone graft procedures. Autografts are preferred but are not necessarily available in sufficient quantities or produce clinically desirable results. Bone replacement materials for maxillofacial, alveolar and mandibular reconstruction are used as an alternative to autografts. Clinically applied techniques include the use of biodegradable membranes for tissue regeneration guided during bone recovery after the implantation procedure. However, despite significant advances in the development of these technologies for the more appropriate three-dimensional nature of composite tissue equivalents, the development of clinically applicable bone replacement materials remains a challenge. ing. At least in part, this challenge is accompanied by difficulties in allowing the repair tissue to ingrow sufficiently within the biodegradable repair material for extended periods of time. As a result, bone construction at the defect site is maintained. Transplantation of such materials at skeletal repair sites usually results in development that is often limited to the surface of the graft, rather than tissue penetrating penetration. However, the latter process appears to be particularly important for successful development and the production of the equivalent of general purpose tissue for maxillofacial and periodontal applications.

  In view of the shortcomings of currently approved synthetic products, including deficiencies in resorption, inclusion of animal or marine-derived components, and poor operational properties, the aim is to produce bone repair materials, Bone repair material behaves both biologically and biomechanically and is more like a maxillofacial bone and a mandible.

Bioceramic fillers have been used to provide the mechanical strength and structural integrity of bone remodeling materials. For example, Ca ₁₀ (PO ₄ ) ₆ (OH) ₂ (hydroxyapatite (HA)) is a bioceramic material that is widely used for bone repair. This is because hydroxyapatite closely resembles the natural tooth and bone crystal structure. For example, US Pat. No. 5,425,769 to Snyders describes an artificial bone replacement composition consisting of fibrous collagen in a calcium sulfate matrix.

  Nevertheless, conventional materials lack composition purity and homogeneity and therefore cannot provide the desired structural integrity. Attempts to solve the problem only give limited success. For example, Jarcho et al., Science 11, 2027-2035 (1976) describes a process for forming dense polycrystalline HA, which is "substantially stronger than other HA materials" and Induces an “excellent biological response when implanted in bone”. A precipitation method was used and material with an average particle size of 150-700 nm was recovered. However, Jarcho et al. (1976) reported significant grain growth during sintering even at a low volume fraction of pores and a firing temperature of 1000 ° C. Jarcho et al. (1976) achieved 99% density in some cases, but used techniques that could be impractical to form the desired morphology. Akao et al. Mater. Sci. 16, 809-812 (1981), reported the compression deflection and dynamic torsional strength of polycrystalline HA sintered at 1300 ° C. for 3 hours, and the mechanical properties of the product, cortical bone, dentin, And compared with the mechanical properties of enamel. The compressive strength of the sintered HA was 3 to 6 times the compressive strength of cortical bone. Hench, J. et al. Am. Ceram. Soc. (1991), HA is used as a coating on a porous metal surface for the fixation of orthopedic prostheses, in particular the HA powder in the pores of a porous coated metal implant It was reported to significantly affect the rate and vitality of bone ingrowth into Many researchers have reported exploring this technique using plasma spray coating of this generally preferred implant. However, Hench (1991) reported that long-term animal experiments and clinical trials of load-resistant dental and orthopedic prostheses suggest that the HA coating can degrade or tear into thin layers. Therefore, application as a filler for bone regeneration is limited.

  Conventional biocompatible materials are difficult to assemble while maintaining their mechanical and structural integrity. For example, HA is difficult to sinter. Thus, high density HA structures for dental implants and low wear orthopedic applications are typically obtained by high temperature and / or high pressure sintering with glass sintering aids, This glass sintering aid frequently induces decomposition into unfavorable phases, poor mechanical stability, and poor chemical resistance to physiological conditions.

  Hydroxyapatite is used in many applications. For example, US Pat. No. 6,241,771 describes a resorbable in-vivo fusion device for use in spinal fixation. The device is composed of 25-100% bioresorbable or resorbable material and neutralizing compound or buffer, and the device is hydroxyapatite. Cambridge Scientific, Inc. EPA 99942186.0 describes a biocompatible tissue graft, which is formed from a solid biocompatible material formed into a suitable form, and the graft is porous on top of it. Formed from a biocompatible biodegradable polymer with a coating agent (where cells capable of regenerating self tissue are seeded on the surface of this polymer coating. EPA by Cambridge Scientific, Inc. 99966346.1 is perforated for implantation by filling the perforations in a perforated and partially desalted cortical bone allograft with a biodegradable porous polymer matrix and partially Describes a method for producing a desalted cortical bone allograft.

  Accordingly, it is an object of the present invention to provide bioresorbable compositions or scaffolds for fillers that increase structural integrity for bone reconstitution.

  Another object of the present invention is to provide bioresorbable compositions or scaffolds for fillers that increase structural integrity for bone grafting or bone tissue regeneration.

  Another object of the present invention is to provide bioresorbable compositions or scaffolds for fillers that increase structural integrity for periodontal, alveolar or maxillary bone regeneration.

(Summary of the Invention)
Disclosed are bioresorbable osteoconductive compositions for bone remodeling and methods of using such compositions. The composition comprises a bioresorbable polymer, a microbiocompatible filler and a nanobiocompatible filler, and preferably a pore forming material. The bioresorbable polymer can be electronically unsaturated and can be crosslinked with a crosslinking monomer. The micro and nano fillers can be any biocompatible material (eg, biocompatible metals, calcium carbonate, biocompatible synthetic materials, carbon, or bioceramics such as hydroxyapatite (“HA”)). possible. The pore-forming material can be a blowing agent (eg, carbonic acid and acid (eg, sodium bicarbonate and acid)), and the acid can be citric acid. In one preferred embodiment, the bioresorbable polymer is polypropylene glycol-fumaric acid. In another preferred embodiment, the monomer is vinyl pyrrolidone. In an even more preferred embodiment, the microfiller and nanofiller are HA.

  The composition can be used for bone tissue regeneration (eg, for mouth reconstruction). In general, mouth reconstruction is associated with the repair of mandibular or maxillofacial defects, maxillofacial bone grafts (eg, sinus augmentation and onlay transplants), or underline stretch. In another embodiment, the mouth reconstruction is periodontal, alveolar or maxillary bone regeneration. In a particularly preferred embodiment, the mouth reconstruction is a tooth replacement. The composition may have other uses (eg, spinal segment repair, skull defect repair) and may be used as a bone graft extender. Bone repair or implantation methods generally include the steps of first producing a suitable template formed from the composition and then implanting the article into a mammal in need of bone repair.

(Detailed description of the invention)
(I. Bioresorbable polymer composition with micro and nano fillers)
(A. Bioresorbable polymer composition)
The polymer must be non-toxic, biodegradable and / or bioresorbable (ie, the degradation product is used by the human body via existing biochemical pathways, or Otherwise it is removed from there). Preferred biocompatible polymers are polyesters or other hydrolyzable polymers. The polyester may be chemically crosslinkable (ie, carry functional groups and allow the polyester polymer chain to react with a crosslinker that is reactive with the functional groups). Suitable polyester materials include polyesters formed from biocompatible dicarboxylic acids and tricarboxylic acids or their ester-forming derivatives (eg, acid chlorides or acid anhydrides) and divalent or polyvalent C ₂ -C _6. Examples include alcohol. The functional groups in the polyester that enable the polyester cross-linking can be derived from the alcohol or acid monomer component of the polyester.

Representative carboxylic acids for polyester formation include Krebs cycle intermediates such as citric acid, isocitric acid, cis-aconitic acid, α-ketoglutanic acid, succinic acid, malic acid, oxaloacetic acid and fumaric acid. Many such carboxylic acids have additional functionality that can allow cross-linking, and thus to process bioresorbable compositions from paste-like moldable masses for hardened cement conditions. Fumaric acid is a preferred acid for forming polyesters, which is a dicarboxylic acid having a free radical reactive double bond that is well adapted to the free radical that induces the crosslinking reaction. there. useful exemplary _C 2 -C ₆ alkyl alcohol or alkylene to form a polyester (aklylene) Arco Are ethylene glycol, 2-butene-1,4-diol, 2-methyl-2-butene-1,4-diol, 1,3-propylene glycol, 1,2-propylene glycol, glycerin, 1,3- Butanediol, 1,2-butanediol, 4-methyl-1,2-butanediol, 2-methyl-1,3-propanediol, 4-methyl-1,2-pentanediol, cyclohexene-3,4-diol In a preferred embodiment, the polyester component of the bioresorbable composition is poly (propylene glycol fumaric acid) (PPF) formed by a condensation (esterification) reaction of propylene glycolic acid and fumaric acid. is there.

  PPF is advantageous. Because PPF possesses two chemical properties that are important for the function of biodegradable bone cement. The first is the ease by which PPF can be degraded in vivo to the original fumaric acid and propylene glycol subunits. Both fumaric acid and propylene glycol are non-toxic and well tolerated in vivo. As a Krebs cycle intermediate, fumaric acid plays an essential role in the process by which food is converted to energy. Propylene glycol is used in the food industry as a food additive and can be metabolized in the body or excreted. A second important property is that each subunit of the PPF prepolymer has an activated site of unsaturation, through which the polyester can be used with crosslinkers that derive various olefinic free radicals. Can be cross-linked.

  This polyester can be crosslinked during curing. When the reactive chemical functional group in the polyester is a carbon-carbon double bond (eg, preferably in the PPF polyester component), a typical cross-linking agent is N-vinyl pyrrolidone (VP). , Methylmethacrylic acid (MMA), and olefin-like crosslinkers. A preferred cross-linking agent is MMA, which exists as a clear liquid at room temperature. MMA is particularly suitable for cross-linking of PPF to induce free radicals according to a preferred embodiment of the present invention.

  Other examples of useful polymers include alpha hydroxy acids (eg, poly (L-butyric acid), poly (DL-butyric acid), poly (DL-butyric acid-co-glycolic acid), and poly (glycolic acid) ), Poly (ε-caprolactone), polyorthoesters, polyanhydrides, polydioxanone, copoly (ether-esters), polyamide polylactones and combinations thereof. These polymers can be obtained or prepared with the molecular weight and molecular weight distribution required for the desired application. Suitable solvent systems for the preparation of these polymers are published in ordinary textbooks and publications. For example, Lang's Handbook of Chemistry, 13th edition, John A. Dean, (eds.), McGraw-Hill Book Co. , New York, 1985. These polymers can be formed into fibers and webs by conventional processing techniques (including melt extrusion and rotary casting) and are commercially available in woven or non-woven form.

  Bioresorbable polymers are known and can be obtained commercially or synthesized using known and published methods. Bioresorbable polymers have been described for various applications, which include controlled release dosage forms and bioresorbable sutures. U.S. Pat. Nos. 3,463,158; 4,080,969; 3,997,512; 4,181,983; 4,481,353; and 4 , 452, 973. Ibay et al., Polym. Mat. Sci. Eng. 53, 505-509 (1985) from vinylpyrrolidone crosslinked to poly (propylene glycol fumaric acid) (PPF) for use as a temporary substitute for soft tissue and / or bone after trauma. The preparation and use of a moldable implant device is described. Resorbable polyglycolic acid sutures have been successfully used for internal fixation of fractures (B. Roed-Peterson, Int. J. Oral. Surg. 3, 133-136 (1974)).

  US Pat. No. 5,522,895 to Mikos described a biodegradable bone template formed with a biodegradable polymer. Useful biodegradable materials include, for example, poly (L-butyric acid), poly (D, L-butyric acid), poly (D, L-butyric acid-co-glycolic acid), poly (glycolic acid), poly (ε -Caprolactone), polyorthoesters, and polyanhydrides, these materials have the ability to become porous. Gerhart et al. Used biodegradable polyesters (US Pat. No. 4,843,112 to Gerhart et al.) That can be chemically cross-linked with a cross-linking agent to form bone cement. Other illustrative examples of bioresorbable or biodegradable polymer applications are US Pat. No. 6,071,982 to Wise et al .; US Pat. No. 4,888,413 to Domb; Yaszemski et al. U.S. Pat. No. 5,733,951 to; and U.S. Pat. No. 4,722,948 to Sanderson.

(B. Pore forming agent)
The porosity of the bone cell carrier promotes bone cell proliferation. For example, US Pat. No. 5,522,895 to Mikos describes a three-dimensional template for the repair and replacement of diseased or damaged bone that is bioresorbable, Provides mechanical support for bone, while providing guidance for bone tissue growth. The template is formed from a biodegradable material in the form of a continuous matrix and a pore-forming component that has a degradation rate that exceeds the degradation rate of the matrix. Differential dissolution or biodegradation provides porosity to this template. Similarly, U.S. Pat. No. 4,722,948 to Sanderson describes a biocompatible polyester resin, a liquid binder that can crosslink with the resin, and a moldable and formable and in vivo curable filling. Bone replacement materials and bone repair materials prepared from the agent were reported. The resulting cured putty will degrade in vivo to provide gaps in the polyester matrix for new tissue growth.

The pores included in the bioresorbable polymer composition can be in any form. In one embodiment, the pores of the bioresorbable composition are biodegradable in the form of fibers and webs having a biodegradation rate that is faster than the biodegradation rate of the bioresorbable polymer. Such materials can be added to the bioresorbable composition. In another embodiment, particles of biocompatible and water-soluble organic or inorganic materials (eg sugar and starch) or organic or inorganic salts (eg NaCl and KCl) are used to form the pores. Can be done. These materials are incorporated into the composition, the composition is solidified, and the particles are extracted using a water extraction technique. Alternatively, the particles can be formed from ammonium carbonate, which is removed by application of reduced pressure or lyophilization. In yet another embodiment, the blowing agent can be used to create pores in the form of foam. In a particularly preferred embodiment, the foam is formed by including a foamable filler that produces CO ₂ when the material is cured. In the most preferred embodiment, citric acid and carbonate or bicarbonate (eg, sodium bicarbonate) are used to form the blowing agent.

The pore size can be controlled by changing the ratio of blowing agent to polymer material and changing the particle size of the blowing agent. A polymer matrix or foam having a pore size of 100-300 microns is desirable to facilitate osteoblast migration. In one embodiment (where sodium bicarbonate (SB) and citric acid (CA) are used as the blowing agent, the foam porosity design is controlled by controlling the SB and CA content, as well as the foam filler. It can be achieved by controlling the size of the SB / CA particles used in the reaction of CA / SB with water to produce carbon dioxide, a blowing agent responsible for foam formation and foam expansion. Requires a molar ratio of CA: SB in the range of 0.2: 1 to 1: 5 (preferably 1: 3) The moles of CO ₂ (which can be produced per gram of material) Depending on the CA / SB filling during foaming, for example, a 0.15% CA / SB filling results in a 25% expansion at 37 ° C. and 1 atm, based on the above stoichiometry.

(C. Micro filler and nano filler)
Microcrystals and / or nanocrystals or nanosynthetic materials are incorporated into the polymer material. Any biocompatible micromaterial or nanomaterial can be used as a filler in the compositions and / or scaffolds disclosed herein. Exemplary micromaterials or nanomaterials are biocompatible metals, calcium carbonate, carbon, biocompatible synthetic polymeric materials, and bioceramics. In one embodiment, the micro or nanomaterial is a bioceramic material such as HA. In another embodiment, the micro or nanomaterial is nickel (Ni), titanium (Ti), aluminum (Ai), gold (Au), platinum (Pt), iron (Fe), silver (Ag), copper (Cu ) A biocompatible metal. By designing materials from the level of clusters, the assembly of crystals below 10 nm can be blocked, and this method can provide unique size dependent properties (eg, quantity limiting effects and superparamagnetism). Methods for foaming various nanoparticulate materials are well known. Various nanocrystalline ceramics for structural applications have been strictly studied in the 1990s. Siegel is “Recent progression in nanophase materials” in B Minerals, Metals and Materials (Suryanarayana et al. (1996)), and many methods exist for the synthesis of nanostructured materials. Synthesis and processing methods to create tailored nanostructures, especially adhesive surface coatings or sufficiently strong, while mechanical or physical vapor deposition, gas condensation, chemical precipitation, aerosol reactions, and biological templates Nanophase metal notes and ceramic notes were considered, requiring only techniques that allow careful control of the surface and interfacial chemistry that can lead to the bulked material. In the case of normal soft metals, reducing the particle size of the metal below the critical length dimension (less than about 50 nm) due to the provision of dislocations in the metal increases the strength of the metal. Metal, intermetallic, and ceramic clusters have been hardened to form ultrafine polycrystals with significantly different and improved mechanical properties over conventional coarse grain counterparts. It is noted that nanophase copper and palladium (assembled from clusters with diameters in the range of 5-7 nm) have a hardness and yield strength that is up to 500% greater than conventionally produced metals. It is also noted that ceramics and conventional brittle intermetallics can be rendered ductile by being synthesized from clusters having a size of less than 15 nm, which is the result of the ultrafine particles being "granular sliding" Resulting from increased ease of sliding on each other. US Pat. No. 5,130,277 to Zettle et al. Describes a method and apparatus for making nanotubes and nanoparticles. A method of forming nano-sized alpha alumna powder from boehmite gel that is doped with a barrier-forming material such as silicon dioxide, then dried, fired, and ground to form a powder is described in US Reported by patent 6,048,577.

  Micro or nanomaterials generally have a particle size in the range of less than 1000 microns, preferably less than 100 microns, even more preferably less than 10 microns, or in the range of nanoparticles. When hydroxyapatite is used to demonstrate microparticles and nanoparticles, there is further bone formation with implants containing nanoparticle HA. Nano-HA is more homogeneous and more pure than conventional HA. Nano HA also has better mechanical properties.

(D. Biologically active material)
Biologically active materials include pharmaceutically active materials and cells that can be incorporated into the composition. Pharmaceutically active materials, such as therapeutic agents such as growth factors and / or other drugs or agents, can be used to enhance bone regeneration and / or tissue adhesion (Lowenguth and Brieden, 1993, Periodontology 2000). , 1: 54-68). There are numerous examples of using growth factors or other pharmaceutically active materials during bone regeneration. Illustrative examples are US Pat. Nos. 4,861,757, 5,019,559, and 5,124,316 to Antonaides et al. Using purified growth factors; US Pat. No. 5,149,691 to Rutherford using growth factor in combination with dexamethasone to enhance mitogenic effect; US patent to Terranova et al. Using mineral loss of root surface. And US Pat. No. 4,916,707, which uses a periodontal boundary such as a membrane. In another example, the periodontal borders are also chemotherapeutic agents (eg, tissue regenerative agents such as growth factors, antibiotic materials, and anti-inflammatory agents to promote periodontal healing and regeneration). Designed so that it can be used for controlled delivery.

  The therapeutic agents can be incorporated for sustained release of these agents in situ. For example, the drug can be incorporated into a polymer matrix or pore-forming material and released slowly as the matrix is degraded. Growth factors, particularly platelet-derived growth factor (PDGF) and insulin-like growth factor (IGF-I), are known to stimulate mitotic, chemotactic and proliferative (differentiated) cellular responses. . Preferred pharmaceutically active materials are materials that increase bone regeneration and / or tissue adhesion. Illustrative examples include growth factors, antibiotic materials, immunostimulants, and immunosuppressants. In one embodiment, the pharmaceutically active material is a bone repair material such as BMP. In another embodiment, the pharmaceutically active material is a growth factor such as an agent that promotes the production of FGF or connective tissue.

  Cells can also be taken up on or in the matrix. Osteocytes can grow in synthetic polymer and natural matrices (Uchida et al., Acta. Orthopedica Scan. 59, 29-33 (1988)). Bone cells obtained from future recipients can be expanded in vitro by manipulating their use in bone repair. For example, Breibert et al. Reconst. Surg. 101 (3), 567-574 (1998), demonstrated the possibility of using periosteal cells for bone repair of calvarial defects by tissue manipulation. Ishaug et al. Biomed. Mater. Res. 28, 1445-1453 (1994); Biotechnol. Bioengin. 50, 443-451 (1996) demonstrated that osteoblasts can proliferate in vitro and migrate through a polymerized scaffold. Ishaug-Riley et al. Then demonstrated in Biomaterials 19, 1405-1412 (1998) that the function of osteoblasts on synthetic materials is equivalent to that of non-degradable orthopedic materials. The seeded resorbable scaffold is placed on new bone when implanted in a site within the body.

  In one embodiment, the composition or scaffold is a PPF-based composition or scaffold where there is sufficient hydrophilicity for cell adhesion or proliferation. When used with the above fillers, the PPF-based composition or scaffold provides a clear porosity for cell migration, creates a rich surface area for neovascularization, and Provide sufficient dimensional stability to support the rebuilding process.

II. Preparation of bioresorbable compositions with micro- or nano-fillers
Preparation of the bioresorbable compositions disclosed herein typically involves forming a formable synthetic cement mass that is cured during processing (ie, completion of the cross-linking reaction). Combining the polymer and the crosslinker with a substantially uniform mixture. The number average molecular weight [M (n)] and molecular weight distribution [MWD] number of the polymer is the number of molecular weights and molecular weight distributions such that the polymer and crosslinker can be combined to form a substantially uniform mixture. Should be. Preferably, the crosslinker is a liquid and the polymer is substantially soluble or miscible in the crosslinker. Alternatively, the crosslinker can be a solid that can be dissolved in a liquid low molecular weight polymer, or can be a liquid miscible therewith. Under ideal circumstances, this cross-linking reaction results in a uniform (uniformly cross-linked) polymer / particle synthetic cement.

  In a preferred embodiment, the poly (polyethylene glycol fumaric acid) (PPF) is sufficient for the reaction to initiate the crosslinking of the polyester to the level necessary to form a hard crosslinked PPF polymer matrix for the mixed particulate calcium salt. Combined with a large amount of methylmethacryl. Preferred MWDs for PPF are in the range of about 500 to about 1200 M (n) and about 1500 to about 4200 M (w). In a preferred embodiment, the liquid polymer phase of the bioresorbable composition is about 80 to about 95% by weight PPF and about 5 to about 20% by weight MMA monomer. The optimal weight percentage for mechanical strength is about 85 percent PPF and about 15% MMA. The MMA monomer is typically stabilized to prevent premature polymerization (ie prior to mixing with PPF) with several parts per million hydroquinone.

  It is important that the ratio of polymer composition and crosslinker is controlled. For example, if too much MMA monomer is added, the MMA molecule will polymerize itself without being blocked by the PPF chain. The result is a material that behaves like conventional PMMA bone cement and does not biodegrade. If little VP monomer is added, the PPF polymer chains will not crosslink efficiently and the cement will not cure to form a sufficiently hard matrix. However, it is known to those skilled in the art how many crosslinkers are used for a particular crosslinker, and this can easily be determined without undue experimentation.

  The VP-PPF crosslinking reaction proceeds by a free radical propagating polymer reaction. This crosslinking reaction is thus accelerated by the addition of a free radical initiator during the run. One suitable free radical initiator for this step is benzoyl peroxide. Other peroxides (eg, t-butyl hydroperoxide and methyl ethyl ketone peroxide) and other free radical initiators (eg, t-butyl perbenzoate) are also suitable free radical initiators for this process. .

  A catalytic amount (less than 1% by weight) of dimethyltoluidine (DMT) is typically added to accelerate the formation of free radicals at room temperature. Accordingly, the rate of crosslinking (ie, curing or hardening time of the bioresorbable composition) can be adjusted by controlling the amount of DMT added to the PPF / MMA mixture. The cure rate is adjusted so that the bioresorbable composition is cured in a time ranging substantially from less than 1 minute to 24 hours or more. The preferred cure time will of course depend on the time that is most practicable for surgical purposes. This cure time is sufficient to allow the surgeon's time to work with the bioresorbable composition to shape or apply the bioresorbable composition to a suitable surface. Should be length. At the same time, this cure rate should be fast enough to achieve, for example, stabilization of the implant within a short time after the surgical procedure. Polymer time or coagulation time for bone implant fixation is typically in the range of about 5 minutes to about 20 minutes, and preferably about 10 minutes.

  A number of procedures can be used to create porosity. In one embodiment (where water-soluble organic or inorganic salt particles are used to create the pores), the particles are leached or otherwise removed from the matrix and highly porous. A polymer matrix having properties remains. In another embodiment (wherein the polymer fibers or webs are dispersed within the formed polymer matrix), the dispersed fibers and the surrounding matrix possess different degradation rates. However, the fibers are broken down at a faster rate than the matrix, whereby the fibers are removed from the template and create a highly porous polymer template. In a preferred embodiment (where the blowing agent is, for example, SB / CA), the pores can be created in the form of a foam upon exposure to water.

III. Use of compositions and scaffolds for bone remodeling
The composition can be used as a scaffold or fixture for the regeneration of any type of bone or tissue. In one embodiment, the composition can be used as a scaffold in periodontal tissue regeneration (eg, soft tissue regeneration, cementum regeneration, or bone regeneration). In another embodiment, the composition can be used as a fixture for bone regeneration (regeneration of spinal segments or repair of skull defects). In yet another embodiment, the composition can be used as a material for bone repair (ie, to not deploy bone but to promote healing). In one preferred embodiment, the composition can be used as a bone graft extender to increase new bone formation when mixed with allograft, autograft, or xenograft material. .

  Bone repair or regeneration can be any type of bone repair, particularly mouth reconstruction, spinal segment repair, bone graft extension. In one particular embodiment, the bone repair is either periodontal regeneration, alveolar regeneration, or maxillary bone regeneration. In one particular embodiment, the bone repair is a tooth repositioning.

  Alternatively, methods of using the bioresorbable composition include 1) ex vivo assembly of a suitable template in the desired form for the desired application, then 2) this template is applied Implanting to the appropriate site can be included, where the template applies to the form of the new material being formed.

  The compositions described herein are useful in the following applications, alone or in combination with other materials (eg, bone grafts).

(A. Clinical support for mandible and maxillofacial defects)
(Bone transplantation)
The bone grafting procedure was almost an indispensable part for implant reconstruction. In many instances, potential implant sites in the maxilla and mandible do not provide sufficient bone volume or bone mass to accommodate a dental implant. This is a result of bone resorption usually caused by the loss of one or more teeth. Bone grafting procedures typically attempted to reestablish bone dimensions, which disappeared due to resorption. Conventional graft materials can be categorized into five categories: a) autografts or endogenous bone grafts, b) allografts or allogenic bone grafts, c) xenografts or xenografts Strip, d) alloplast or dysplastic bone graft, and e) growth factor.

  Autografts or endogenous bone grafts are considered gold standards. The most successful speed in bone grafting has been achieved using autografts. For most transplant purposes limited to oral transplantation, the part of the jaw (ie, the heel or the posterior part of the jaw) can be used as an acceptable donor site. However, sometimes the iliac crest is recovered if there is not enough bone volume available in the mouth. The source of this allograft is usually cadaveric bone, which is available in large quantities. The cadaver bone must undergo many different treatment sequences to neutralize the cadaver bone to the immune response and avoid cross-contamination of host disease. These treatments can include irradiation, lyophilization, acid washing and other chemical treatments. Xenografts or xenograft bone grafts are often of bovine origin. The tissue bank usually selects this transplant material. Because it is possible to extract a large amount of bone with a specific microstructure, this microstructure is an important factor for bone growth compared to bone of human origin. Alloplast or alloplastic bone grafts include any synthetically derived graft material, usually not derived from animal or human origin. In oral transplantation, this usually includes hydroxyapatite or any formulation thereof.

(Sinus increase)
One of the most frequently applied transplantation procedures is sinus augmentation. This procedure is limited to the upper jaw. Sinus vacuolation occurs with age. Once the tooth disappears in that particular area, this makes it difficult to place an intraosseous implant in that area. Because of this particular problem, the implantation method is to make the sinus base completely tall and implant bone under it, thus creating sufficient clearance for one or more dental implants. ,It has been developed.

  A significant volume of bone (5 cc to 10 cc per side) is required to perform a typical sinus augmentation. This bone mass is usually higher than can be recovered from a donor site in the mouth. Thus, allografts, alloplasts or xenografts or combinations (sometimes mixed with small amounts of autografts) are sometimes necessary. However, autografts usually take approximately 4-6 months to mature within the sinus; while allografts, alloplasts or xenografts can take more than 9 months to mature.

  Sinus augmentation and implant placement can sometimes be performed as a single procedure if sufficient bone between the maxillary ridge and the base of the sinus can be utilized to sufficiently stabilize the implant. If insufficient bone can be utilized, this sinus augmentation must be performed first, and then the graft must mature for several months, and this maturation depends on the graft used Depends on the material. Once the implant is mature, the implant can be placed. In this case, as described above, the compositions provided herein can be used in augmenting endogenous bone grafts.

(Onlay transplant)
The onlay implantation procedure is designed to re-establish bone, which has been lost in a particular area due to resorption (again caused by the loss of pre-tooth in that area). ing. Usually, some endogenous bone fragments (usually from the back of the heel or mandible) are glued to the site with the bone defect. The area is then closed, and after a certain healing and maturation period, the bone fragments are eventually incorporated into the host floor and fused tightly, so that at a later stage, the implant is placed in the same area. obtain. If the greater area of resorption needs to be augmented with more endogenous bone fragments, the patient's bone from the iliac crest or tibia is used.

(Uplift expansion)
If the original prosthetic implant cannot be placed because the jaw ridge is so thin, ridge enlargement is used to restore the lost bone dimensions. In this procedure, jaw bone mass is effectively expanded by mechanical methods. A series of bulking agents (which can be round with a continuously increasing diameter cross-section or a D-shaped metal rod) are forced into the implant site. This is accomplished by using a surgical scissors to drive these extenders into the ridge. If done properly, the use of a bulking agent pressurizes the spongy part inside the bone and protrudes out of the external cortex. In this regard, a suitable implant can be immediately placed in the created socket, or a bone graft can be placed in the initially created socket, and the bone graft before the implant is placed. Can mature in a few months.

(Increase)
Newer bone graft materials are being tested for augmentation procedures. However, the use of osteoconductive bone substitutes in this representation is controversial. It has been postulated that its use can lead to prolonged healing times, uneven bone formation, foreign body reaction, particle movement and low bone-implant contact. To eliminate this problem, the combination of osteoinductive protein (recombinant human bone morphogenetic protein-1 (rhOP-1 = bone morphogenetic protein-7)) and natural bovine bone mineral (BioOss) was added to the implant simultaneously. It has been studied in cavernous augmentation using placement. Terheyden et al. (1999) augmented the maxillary sinus floor in 5 minipigs with 3 mL BioOss containing 420 micrograms rhOP-1 on this test side and 3 mL single BioOss on the control site. It was. Upon augmentation, a titanium implant (ITI) was inserted from the lateral tail direction. After 6 months of healing, the percentage of bone-implant contact was higher than 42% in this increased group. It was concluded that the recombinant growth factor can be delivered by natural bone mineral.

  Distraction osteogenesis has been proposed for both closure of fistulas in patients with extensive alveolar clefts and fissures and reconstruction of maxillary alveolar defects in traumatic patients. The goal is to adhere to the gingiva for the creation of new alveolar bone segments and the reconstruction of complete proximal and maxillary alveolar defects of wide alveolar cleft / fistula. This procedure has recently been performed for one patient with traumatic maxillary alveolar defect and 10 patients with ruptured lip and palate on one or both sides (with altered alveolar cleft / fistula) (Liou et al., Plast. Reconstr. Surg., 105 (4), 1262-72, 2000). Interdental and maxillary osteotomy was performed on one side of the dental arch due to tears or defects. After a 3 day incubation period, the distal segment of the osteotomized dental arch was then dispersed and moved in the cleft or defect direction using an intraoral distribution device. The alveoli and gingiva on both ends of the crack or defect were approached after dispersed bone formation. This method can eliminate the need for extensive alveolar bone grafting.

(Dental implant)
The type of dental implant used often depends on the condition of the maxillofacial bone. This bone thickness and volume determine the type of implant installed. Furthermore, transplantation and reconstruction techniques often require placement of dental implants in the first step. In general, dental implants can be categorized into three main groups: (1) intraosseous implants, (2) subperiosteal implants, and (3) transmyelinated implants.

  The intraosseous implant is surgically inserted into the mandible. The subperiosteal implant is typically the upper part of the mandible but below the rubbery tissue. The important difference is that these implants do not normally penetrate into the jawbone. Transmyelocytic implants are by definition similar to intraosseous implants in that they are surgically inserted into the mandible; however, they differ in their orientation. Intraosseous implants are the most frequently used implants today. Examples of each implant are described below.

(Ramusframe implant)
A branch frame implant is an intraosseous implant. These implants are designed only for the mandible with missing teeth and are surgically inserted into the jawbone in three different areas: the left posterior and right posterior areas of the jaw, and the heel area in front of the mouth . These types of implants are commonly used in severely resorbed mandibular bones that have lost their teeth, and this mandible has sufficient bone height to fit the original implant when fixing the device. Do not provide. These implants are usually indicated when the jaw is resorbed to the point where the subperiosteal implant is no longer sufficient. A further advantage of this type of implant is the stabilization of the lower jaw at three points. Once surgically inserted, a rod running from one side of the jaw to the other is visible in the mouth. The denture can then be glued to this bar.

  Although winged implants are not used frequently, they have a very thin forehead residual bone protuberance (due to resorption) that prevents traditional original implants from being placed or certain critical anatomical features It does not find application in areas where the structure either prevents the placement of conventional implants. Frequently, if a particular area of the jawbone is very thin and undergoes resorption due to loss of teeth, it is recommended to undergo a bone grafting procedure, which reestablishes bone loss and As a result, a conventional original implant can be placed. For such applications, the materials described herein are particularly well adapted.

(Subperiosteal implant)
Of all devices currently in use, the subperiosteal implant is the type of implant that has the longest clinical application. These implants are shaped to ride on the remaining bone protuberance of either the maxilla or the mandible. Subperiosteal implants are used in the fully and partially edentulous maxilla and mandible. However, the best results have been achieved in edentulous mandibular treatment. Indications usually include severely resorbed mandibular bone with missing teeth, which does not provide sufficient bone height to fit the original implant when fixing the device .

(Original implant)
Due to bone integration, these titanium implants have become the most common implants. They are considered standard care in oral transplantation. If sufficient bone can be utilized to fit these implants, these implants can be placed anywhere where the teeth are missing. However, even if the bone volume is insufficient to place the original implants, a reasonable range of bone grafting procedures should be initiated to benefit from these implants. Some newer implants have an outer coating of hydroxyapatite. Other implants have a surface modified by a plasma spraying process, an aderchim process or a beading process. Other variations focus on the shape of the original implant. Some are in the form of screws, others are in the form of cylinders or even cones, or any combination thereof.

  The microparticle or nanoparticle compositions described herein may be used: (1) high density ingrowth of bone cells within the scaffold, (2) integration of ingrowth and surrounding tissue after implantation Can be used to support the reconstruction process by allowing (3) angiogenesis and (4) cosmetic surgery recovery. The method of use of this composition will vary depending on the particular procedure for the particular desired application.

(B. Periodontal tissue regeneration for implant support)
The materials and methods disclosed herein can be used for periodontal regeneration. Periodontal regeneration may be a type of soft tissue, cement or alveolar bone healing characteristic of undesired periodontal tissue dissection and construction. In general, periodontal regeneration involves inserting a subperiosteal implant into a mammalian dental ridge. Alternatively, a therapeutically effective amount of growth factor can be used with the insertion of a subperiosteal implant. The inserted periodontal regeneration system can then be shaped to form a periodontal boundary when used for the treatment of periodontal disease around the root. This boundary can be positioned to create and maintain a gap for regeneration between the gingival tissue and the root surface. Finally, the wound is closed to allow periodontal regeneration.

  Insertion of a subperiosteal implant generally includes one or more of the following steps: 1) making a side incision in the tissue covering the bone ridge (this incision may extend beyond the bone ridge). 2) tunneling the tissue both proximally and distally from the incision so that the tissue separates from the bone ridge; the proximal and distal distances of the tunneling tissue are arranged Equal to the length of the periodontal regeneration system, 3) injecting the periodontal regeneration system under the tissue in a proximal orientation and a distal orientation, respectively, 4) implanting the bone Fixing the two parts of the implant together by molding into a desired shape under the tissue, 5) suturing the incision, and 6) a post-periosteal implant system for post mounting The Been, the step of inserting on the preassembled subperiosteal implant. Those skilled in the art will determine whether one or more of the preceding steps are required for a particular periodontal regeneration and / or which of the preceding steps are required for a particular periodontal regeneration Can be determined.

  Alternatively, a therapeutically effective amount of growth factor can be applied with the periodontal regeneration system. The growth factor can be, but is not limited to, one or more of the following: platelet derived growth factor in the form of two beta chains (PDGF-BB), platelet in the form of alpha and beta chains Derived growth factor (PDGF-AB), IGF-I; and TGF-β or their precursors in either DNA or mRNA form.

  The periodontal disease can be any wound of periodontal disease that requires bone or tissue repair or regeneration. For example, the wound can damage mammalian bone, periodontal tissue, connective tissue, or ligaments. In one embodiment, the defect is one of class III bifurcation lesions or other periodontal tissue defects resulting from periodontal disease or other destructive or trauma processes to periodontal tissue. is there.

  Methods of using the compositions disclosed herein are further understood as a reference to the following non-limiting examples.

Example 1. Nano HA particles or micron HA particles to augment PPF bone grafts
Bioactivity of bioresorbable bone graft substitutes made from unsaturated polyester, poly (propylene fumaric acid) augmented with nanohydroxyapatite, biocompatibility of implants placed in rat tibia defects and bone Analysis was performed by assessing osteointegration. Three groups, including eight animals in one group, were evaluated by grouting bone graft replacements into 3-mm holes made in the metaphysis of the rat anterior central tibia. Two different formulations were used, varying depending on the type of hydroxyapatite; group 1: nanohydroxyapatite, group 2: micron hydroxyapatite, group 3: HA only control. Three groups of animals each were sacrificed 3 weeks after surgery in 8 groups. Histological analysis showed excellent biocompatibility and bone integration of bone graft replacements when using nanohydroxyapatite. After 3 weeks there was more reactive new bone formation in this group when compared to the micron hydroxyapatite group. This control group showed incomplete closure of this defect. This study showed that nanohydroxyapatite improves the biological activity of bone implants and repair materials. The model scaffold used in this study (poly (propylene fumaric acid)) provided a bone conduction pathway through which bone grew faster.

(Materials and methods)
(Materials and formulations)
A typical formulation used for this study is shown in Table 1. PPF (Mw is approximately 5,000 by GPC), Gresser et al., J. MoI. Biomed. Mater. Res. 1995, 29: 1241-1247 and Gresser et al., Bone cement Part 1: Biopolymer for avulsive maxilofial repair, Human Biomaterials Applications, Wise DL, Greser JD, TransY. , Totowa, New Jersey, 1996, 169-187 (1996), and synthesized from equimolar fumaric acid and propylene glycol in the presence of p-toluenesulfonic acid. 1-vinyl-2-pyrrolidinone (VP), benzoyl peroxide (BP), hydroquinone (HQ), and NN-dimethyl-p-toluidine (DMPT) were purchased from Aldrich (USA) and were widely accepted. Used as a thing. Sodium bicarbonate (SB) and citric acid (CA) were purchased from Fisher Scientific (USA).

  To form a viscous putty-like paste that results in a cross-linked polymer, a liquid component consisting of VP, accelerator DMPT, and sterile water (Part II) is combined with PPF, HA, SB, BP initiator, and CA. To a dry powder mixture consisting of (Part I). The accelerator (DMPT) gave a working time of about 90 seconds at a concentration of 0.03% w / w. The reaction between CA / SB and water produces carbon dioxide, a spray, responsible for foam formation and expansion. The stoichiometry follows the method of Bondre, Tissue Engineering, 6 (3): 217-227 (2000) and has a 1: 3 molar ratio of CA: SB with a CA: SB weight ratio of 1.00: 1.31. Request.

１００〜３００ミクロンの孔サイズを有するＰＰＦ泡沫は、骨細胞の内殖のために望ましいようであった。ＣＡ／ＳＢと水との反応は、孔形成および拡大を担う二酸化炭素、吹き付け剤を産生する。化学量論は、１．００：１．３１のＣＡ：ＳＢ重量比を有する１：３のモル比のＣＡ：ＳＢを要求する。ＣＯ_２のモル比（材料１ｇあたりに産生され得る）は、セメントを泡立てる際のＣＡ／ＳＢの充填に依存する。０．１５％のＣＡ／ＳＢの充填は、上記の化学量論に基づいて、３７℃および１ａｔｍで２５％の伸張を生じる。この吹き付け剤に加えて、ＰＰＦ処方物は、Ｂｏｎｄｒｅ（２０００）による記載される技術を使用して、骨伝導性のＨＡ充填剤の存在下で、ビニルピロリジノンを使用して架橋した。

PPF foam with a pore size of 100-300 microns appeared to be desirable for bone cell ingrowth. The reaction between CA / SB and water produces carbon dioxide, a spray, responsible for pore formation and expansion. The stoichiometry requires a 1: 3 molar ratio of CA: SB with a CA: SB weight ratio of 1.00: 1.31. The molar ratio of CO ₂ (which can be produced per gram of material) depends on the CA / SB loading when foaming the cement. A loading of 0.15% CA / SB results in 25% elongation at 37 ° C. and 1 atm, based on the above stoichiometry. In addition to this spray, the PPF formulation was crosslinked using vinylpyrrolidinone in the presence of osteoconductive HA filler using the technique described by Bondre (2000).

The hydroxyapatite used in this study was: nano HA (Group 1, average particle size = 40 nanometers as produced and characterized by professor Ying of MIT) and sintered 10 μm HA, Spherical micron HA (Group 2, CAM Implants, commercially available from The Netherlands, average particle size = 26 microns) [Panchula and Ying, in Nanostructured Materials: Science and Technology, ed by GM Chow and Nw
, 1998), 319-333); Sun and Ying Nature 389: 704-706 (1997); Ying, Designer materials through nanoprocessing of Frontier of Engineering, 1996, National Ac., National 23. ~ 27; Ying et al. Am. Ceram. Soc. 76 (10): 2561-2570 (1993); Ying et al., Angew. Chem. Int. Ed. 38 (1): 56-77 (1999); Zhang et al., Chem. Mater. 11 (7): 1659-1665 (1999); Zhang et al., Chem. Commun. 1103-1104 (1999)]. This nano-HA was compared to micron HA and an empty defect (Group 3), which was left for spontaneous healing. All HA preparations used in this study were subject to X-ray diffraction (XRD) to study crystal purity and size, photoacoustic Fourier transform infrared (PA-FTIR) spectroscopy to demonstrate molecular structure, and Characterized using a transmission electron microscope (TEM) to determine particle size and porosity.

(In vivo animal research and group design)
The three groups were divided into Gerhart et al., J. et al., Using any of the available HA types (Groups 1 and 2) and an empty control (Group 3). Orthop. Res. 3, 11: 250-255 (1993), used in rat metaphysis model of tibia. NIH guidelines for laboratory animal healing and use (NIH publication, # 85-23, Rev. 1985) were observed. Adult male Sprague Dawley rats weighing approximately 200 g were used as an animal model (Zivic Miller, Zelenople, PA, USA). The animals were anesthetized using intramuscular injection of ketamine hydrochloride (100 mg / kg) and xylazine (5 mg / kg). Rats were also given an intramuscular prophylactic dose of penicillin G (25,000 U / kg), and the surgical site was shaved and Betadine (Povidone iodine) and alcohol (Dura-Prep; 3M Health Care, St. Paul, MD) , USA). A 1.5 cm longitudinal incision was made in the anterior part of the left hind leg and the tibia metaphysis was exposed. A 3 mm hole was made in the anterior midtibial metaphysis of the rat. This formulation (mixed to a viscosity similar to paste or putty before surgery) was implanted using a spatula into the prepared tibial defect site. The PPF-based grout was hardened in situ, and after implantation of the bone grout, the soft tissue and skin were closed in layers with absorbable sutures. A single formulation was transplanted into 8 animals. All animals were sacrificed 3 weeks after surgery.

(Evaluation methods)
Evaluation was performed by high-resolution radiographs taken directly using post-surgery and 3 week intervals from specimen x-ray units (Microfocus 50E6310F / G; Xerox, Rochester, NY, USA). Radiographs were taken with minimal exposure (32 kvp, 2 seconds), and mammography film (Clonex Microvision; Dupont Medical Products, Wilmington, DE, USA), cassette (MR Detail; AGFA Richfield Park, NJ, USA) and screen Mammolay (AGFA) was used. After sacrifice, 10 mm long tibial segments (including sections to be transplanted with bone graft replacements) were collected. This specimen was processed by fixation in 10% buffered formalin for histological analysis. Specimens (including residual bone graft material) were decalcified in EDTA and embedded in paraffin. The longitudinal sections (5 μm thickness) of the total specimen were then sliced and stained with hematoxylin and eosin. In addition, slides were stained using the von Kossa method to demonstrate calcium crystals. Slides were tested for resorption activity and new bone formation at the site of implantation, and inflammatory response.

  Histomorphometric evaluation of new bone formation around different types of grafts, images of serial serial sections of this specimen stained with hematoxylin and eosin, mounted on a Zeiss microscope (IM- 745; PULNiX, Sunnyvale, CA, USA). Images were digitized and analyzed using Image Pro Plus software. For each specimen, the area of newly formed bone around and within the implant was measured. This measurement was normalized to the total area occupied by the implant in the same section. A minimum of 5 sections from different levels of this specimen were included for this analysis. The spacing between adjacent level sections is typically 300 micrometers, which takes into account the approximate absolute volume of the newly formed bone, and this absolute volume is calculated for each bone to be obtained. The sample is given as an average percentage ratio of these volume measurements (mean ± standard deviation). To compare the extent of new bone formation around the graft at the metaphyseal site between the experimental and control groups, a recovery index was determined. The recovery index was defined as the volume ratio of newly formed bone and the volume of all implants based on 8 animals per study group. Thus, absolute capacity is given as an average percentage ratio.

(Statistical analysis)
Differences in the reconstruction index were analyzed for statistical significance by using the ANOVA test. A p-level of 0.05 was considered statistically significant.

(result)
(Control group)
In the control group, there was no new bone formation in the metaphyseal defect made in the rat tibia. At the site of cortical drilling, there was some periodontal bone formation, but the remaining portion of the defect at the tibial metaphysis was mainly filled with bone marrow and adipose tissue.

  In the micron hydroxyapatite group, the implant remained structurally stable and did not collapse. There was no histological evidence of implant degradation or active cell resorption from this recipient site. There are some moderate infiltrates with PMN, which seemed to be consistent with post-operative inflammatory changes. In addition, there was new bone formation, which appeared to fill tightly around the implant 3 weeks after surgery, without excess fibrous or inflammatory tissue. There is osteoclast and osteoblast activity on the surface of the implant, suggesting that the bone around the implant undergoes active remodeling. Although the bone surrounding the implant bone graft undergoes active remodeling, the implant remained structurally intact. Microhydroxyapatite crystals were readily demonstrated using von Kossa staining. It was noted that new bone formation and large numbers of round cells, whose morphological appearance is consistent with osteoblasts, are in close proximity to micron HA crystals.

  In the nanohydroxyapatite group, the implant surface stimulated a more active inflammatory response with infiltration by PMN and macrophages. Furthermore, there appeared to be additional new bone formation around this implant. Similar to group 2, there was also no histological evidence for implant lysis or active cell resorption from the recipient site in group 3. In contrast to the micron hydroxyapatite group, HA crystals were unstainable using the von Kossa technique in the nano hydroxyapatite group. Similarly, in the micron HA group, there were large cells (positioned in the direction of the interface with this implant) having a circular nucleus. Osteoid seemed to be secreted onto this implant material.

  Histomorphometry was measured in the control group (no graft; p <0.002) and the micron HA group (p <0.025) around the different types of grafts used in this study. It was significantly higher in the nanohydroxyapatite group than in Both formulations were equally osteoconductive but covered by newly formed reticular bone, as measured by this implant area, the wider edge of the newly formed bone was nanohydroxyapatite Care was taken around the implant. In addition, additional new bone has been found in these types of implants. As a result, the reconstruction index was higher in the nanohydroxyapatite group when compared to the micron hydroxyapatite group (see Table 2).

ラットにおける本動物研究の最終的な目的は、骨伝導性のグラウト中でナノヒドロキシアパタイトの用途有用性を確立することであり、一方で、異物反応の非存在で生体適合性を証明し、そして骨治療／修復工程に対する効果を評価する。本明細書中に存在するＰＰＦベースの再吸収可能な骨移植片置換は、骨伝導性であることが予測された。なぜなら、このヒドロキシアパタイト充填剤は、類似のモデル評価において首尾よく使用されているからである。

The ultimate goal of this animal study in rats is to establish the utility of nanohydroxyapatite in osteoconductive grout, while demonstrating biocompatibility in the absence of foreign body reaction, and Evaluate the effect on the bone treatment / repair process. The PPF-based resorbable bone graft replacements present herein were predicted to be osteoconductive. This is because this hydroxyapatite filler has been used successfully in similar model evaluations.

  Two types of HA particle sizes were studied in this study: (a) sintered spherical nano-HA (average particle size = 40 nanometers), and (b) sintered spherical micron HA (average particle size = 26 nanometers). The results of this study did not show new bone formation in the negative control, which was not filled with any graft material. At 3 weeks, more reactive new bone formation occurred in the nano HA group than in the micron HA group without completing the defect. These histological observations were supported by histomorphometric measurements of new bone formation that demonstrated a significant increase when PPF was used in combination with nano-HA (Table 2). There is no evidence of implant failure or degradation during the period analyzed in this rodent study (3 weeks). Furthermore, it is clearly shown that these PPF-based bone graft substitutes are overly osteoconductive and exhibit both newly formed reticular bone ingrowth with concurrent neovascularization. certified.

  Based on these observations, the present study showed that the osteoconductive properties of PPF-based bone graft materials containing 13.7% HA can be further improved by the use of nanohydroxyapatite. Rapid bone ingrowth and healing can be facilitated by accelerated bone formation around and within the biodegradable scaffold.

(Example 2: Repair of periodontal defect having PPF containing HA particles)
(Materials and methods)
Poly (propylene fumaric acid) was synthesized by direct esterification of fumaric acid (Fisher Scientific, Inc.) with propylene glycol (Aldrich Chemical Co., Milwaukee, Wis.) As described above. In brief, the reaction is carried out at elevated temperatures in the presence of t-butylhydroquinone (Aldrich Chemical Co., Milwaukee, Wis.) (Inhibitor of natural crosslinking) (Aldrich Chemical Co.). , Milwaukee, WI). The reaction product is dissolved in methylene chloride, filtered to remove unreacted fumaric acid, washed with 20% aqueous methanol to remove unreacted propylene glycol, and type 3A Dry with molecular sieves (EM Science Co.). The polymer was recovered from methylene chloride by precipitation in diethyl ether, redissolved in acetone, dried, filtered, and the acetone removed under reduced pressure. The weight average molecular weight of PPF, 7.8 × 300 mm of ultras fliers gel 10 ³ Angstrom column having (Waters, Model 410, Milford, MA) ( dispersity 2.57 [17,20], it should be 6650 Determined by gel permeation chromatography. Hydroxyapatite (ha, CAM Implants BV, The Netherlands) was made in micro or nano form. N-vinylpyrrolidone (VP) (Aldrich Chemical Co., Milwaukee, Wis.) Was distilled under reduced pressure (93 ° C., 13 MMHg) to remove NaOH inhibitors. VP in solution with n, n-dimethyl-para-toluidine (DMPT) (Aldrich Chemical Co., Milwaukee, Wis.) and Tween 80 (Fisher Scientific, Inc.) forms a viscous tannin colored pate. To the dry powder mixture of PPF and HA. Sodium bicarbonate (Fisher Scientific, Inc.), benzoyl peroxide initiator (Aldrich Chemical Co., Milwaukee, Wis.), And citric acid (6.8 wt% solution) (Fisher Scientific, Inc.), in turn, Added to make a conductive foam network.

This unsaturated PPF polymer can be cross-linked with VP in the presence of blowing agents, sodium bicarbonate (SB) and citric acid (CA), and HA to create an osteoconductive foam network, Can be mixed with the bone graft just prior to filling the defect. The reaction of citric acid and sodium bicarbonate with water produces a spray agent responsible for carbon dioxide, foam formation and foam expansion. A 1% SB / CA charge results in an expansion of about 200% at 37 ° C. and 1 atm based on stoichiometric release of CO ₂ according to the following reaction:

５００ｌｂのロードセルを備える試験機を使用して、そして１ｃｍ／分のクロスヘッドスピードで操作すると、平均強度および係数は、それぞれ、１７．７±２．８ＭＰａおよび３６５．３±７４．９ＭＰａであった。これらの圧縮強度データは、ＣａｒｔｅｒおよびＨａｙｅｓにより報告された値と比較可能であり、彼らは、骨のこの特性を歪み速度および歪み密度の機能として測定した。比較可能な歪み速度で、海綿質の圧縮強度（ｐ＝０．３１ｇ／ｃｍ^３）は、５．０ＭＰａであり、そして皮質骨については、海綿質の圧縮強度（ｐ＝２．０ｇ／ｃｍ^３）は、２００ＭＰａであることに留意した。

Using a test machine with a 500 lb load cell and operating at a crosshead speed of 1 cm / min, the average strength and modulus were 17.7 ± 2.8 MPa and 365.3 ± 74.9 MPa, respectively. . These compressive strength data are comparable to the values reported by Carter and Hayes, who measured this property of bone as a function of strain rate and strain density. At comparable strain rates, the sponge compressive strength (p = 0.31 g / cm ³ ) is 5.0 MPa, and for cortical bone, the cancellous compressive strength (p = 2.0 g / cm ^3). ) Was 200 MPa.

(Decomposition mechanics)
Polymer degradation occurs by hydrolysis of ester bonds, which results in chain scission (Gresser et al., 1995; Peter et al., 1997a, 1997b). Quantitative development of mechanical data in vitro is based on the initial compression value (sponge having a ratio of mechanical disappearance, not exceeding 50% of the initial value in 3 weeks (5 Mpa for its strength and its coefficient Is focused on the desired in vitro results, currently defined as being comparable).

In order to determine the impact of polymer degradation on machine disappearance over time, an in vitro assessment of the biomechanical properties of the scaffold was performed with respect to polymer degradation, and the time sequence of polymer degradation and the mechanical strength of this graft material Were correlated. Samples for modulus testing and strength testing were performed in phosphate buffered saline at 37 ° C. (ASTM Method F1634-95, “Standard Practice for In-Vitro Environmental Conditioning of Polymer Matrix Composites”). And so on) under load to the appropriate physiological conditions and removed at day 4 and at 1 week, 2 weeks, 3 weeks, 6 weeks, 9 weeks and 12 weeks. The protocol for loading samples in vitro was proposed in the FDA Guidance Document for Testing Implant Devices (April 1996) and was previously used for in vitro testing of biopolymer immobilization under load (Trantolo et al., 2000). Based on the application of ASTM Methods F451-99a (specifications for acrylic bone cement) and D5024-95a and D4065-95 (for measurement and reporting of dynamic mechanical properties). When removed from the bath, the sample was maintained in a gauze soaked in saline under conditions of fully hydrated. Dr. Oregon Health Sciences University laboratory. Wilson C.I. Using the standard procedure used by Hayes, the mechanical value of the PPF foam extender was determined in an in vitro evaluation. Automatic data collection was used throughout the study using Lab View ^™ . All experiments were performed using Standard Operating Procedures (SOP's) for each protocol and data analysis as required by the FDA guidelines under the Good Laboratory Practice (GLP) guidelines.

  Analysis of variance was used to detect statistically significant differences in the rate of machine disappearance measured at each study interval. A paired t-test was used to compare micron and nano-HA formulations. In all statistical tests used, a significant level of p <0.05 was selected.

  The formulation selected for the mechanical basis was subjected to morphological characterization by scanning electron microscopy (SEM, AMR-1000, Advanced Metals Research Corp.). Temporal profiles were processed using NIH Scion Image software at 6 time points (t = 0 week, 1 week, 3 weeks, and 6 weeks) using 6 samples of each selection. And developed using the analyzed data.

(Evaluation procedure)
PPF-based grouts enhanced with either nano-HA or micron-HA were screened and compared to controls. Rats were divided into four groups, each of which had a micron HA PPF implant (Group 1), a nano HA PPF implant selected for Task 2 mechanical results (Group 2), or mineral shedding Any of the bones that were made (Group 3) were implanted. The fourth group (Group 4) was a sham control (ie, an unfilled defect) to assess the spontaneous healing of this defect model. 45-day-old Sprague-Dawley rats (Zivic Miller, Zelenople, PA, USA) were divided into 4 groups with 36 animals per group.

  For tooth extraction, the animals were weighed and then anesthetized by a single intraperitoneal injection of bentobarbital sodium (Nembutal) at 50 mg / kg body weight. A sharp probe was used for tooth extraction. All right maxillary and mandibular molars were removed under a dissecting microscope (Dentiscope, Johnson & Johnson, E. Windson, NJ). The periodontal ligament was then relaxed from the tooth neck of each tooth by placing the tip of the probe around each tooth and gradually separating the tissue. The short needle was then placed between the molars and the molars removed.

  In the experimental group of animals, the mandibular region was packed with a different PPF-based implant material, or a mineral shed human bone matrix, commercially available as Grafton Putty® from Musculoskeleton Transfoundation Foundation. After the material was placed in the alveolar together with the amalgam carrier, the alveoli were carefully packed with a small amount of amalgam condenser. A nodule suture was attached to close the extraction site and maintain the implant material in place. Animals were given soft rat chow and water as appropriate.

  Twelve animals from each group were sacrificed at 3, 6, and 16 weeks after surgery. Two weeks prior to sacrifice, animals were injected with fluorescent dyes for histological evaluation as described below. A total of 144 rats were included.

Assessments are performed by using radiographic, histological and histomorphological techniques and these techniques are described below:
(Radiograph)
Standard radiographs (0.3 s, 80 kVp) were taken on the first day, and at 2, 4, and 16 weeks after tooth extraction. Two radiographs were made for both the right and left sides of the mandible. In order to normalize the angle of the X-ray beam, an acrylic cone guide was fixed to the skull measuring instrument. To measure alveolar height as described by Nishimura et al. (1987), radiographs were scanned and analyzed digitally using Image Pro Plus software. As early work has demonstrated, films made by this technique are substantially superimposable.

(Histological experiment)
Animals were sacrificed at 3, 6, and 16 weeks after tooth extraction and PPF implantation. The lower jaw was removed directly, skinned and fixed in 10% formalin for more than 24 hours. The specimen was shaped to include only the basal area of the molar and then placed in ethylenediaminetetraacetic acid (EDTA) until the specimen was fully decalcified. Specimens were cut along their length in the near-centrifugation direction and embedded in paraffin. Sections 5 μm thick were made and stained with hematoxylin and eosin.

(Histomorphometry)
Tissue morphometry was then performed on all specimens. For this purpose, a CCD video camera system (TM-745; PULNiX, Sunnyvale, CA, USA) attached to a Zeiss microscope was used. Images were digitized and analyzed using image software (ImagePro Plus). For each calvarial tissue specimen, the area of the defect that was replaced by newly formed bone (ie, the implant) divided by the total area of this defect was determined on serial tissue sections. The spacing between adjacent level sections was typically 300 micrometers. A minimum of 10 sections were included for this analysis. Data were presented as average percentage ratio (mean standard deviation).

(Statistical analysis)
A Mann-Whitney test was performed to determine a statistical analysis of the change in alveolar height between the untreated control group and the implant treated group. After extraction in the untreated control group and the implant-treated group, a Mann-Whitney test was performed for net weight gain to test the null hypothesis that the net weight gain was the same in both groups. The resorption pattern for each animal was determined to fit a resorption curve that fits the quadratic equation y = ax ² + c. Multiple linear regression analysis was performed for each animal used in this study. The coefficient of determination (R ² ) showed a good fit. R ² is a ratio and is a value in the range of 0 to 1; the closer the value of R ² is to 1, the better the individual data fit with the reabsorption model. The third analysis is a study of the change in results at a particular time by using Kruskal-Wallus analysis. This non-parametric one-way analysis of variance was used to determine whether reabsorption rates differ between groups at specific times during the study (3 weeks, 6 weeks and 16 weeks after tooth extraction). Applied to the untreated control group and the implant treated group. Finally, a t-test was used to measure the change in striated height between 3 and 16 weeks after tooth extraction. This test was applied separately to the untreated control group and the implant treated group.

  This formula was evaluated by implanting a graft in the mandibular region of the rat.

(result)
Crosslinking of porous scaffold with unsaturated PPF polymer and vinyl monomer, vinyl pyrrolidone (VP) in the presence of foaming filler, sodium bicarbonate and citric acid, and osteoconductive filler (HA) Formed by. Upon mixing, the mixture cures by cross-linking of the PPF with simultaneous monomer and CO ₂ yielding a porous scaffold that is degradable by hydrolysis. The use of HA as part of the filler supports the scaffold's osteoconductivity (Saito, Biomaterials 15, 156-160 (1994)), while the pores created by CO ₂ are in situ cells. It provides a porous region for adhesion and growth (Bondre et al., 2000), and the hydrophilicity of this polymerized support facilitates cell migration (Lewandrowski et al., 1999).

  PPF scaffolds (without HA particles) were evaluated for morphological, mechanical and surface properties in vitro and in vivo using a rat tibial defect model (Gerhart et al., 1989). The scaffolds designed for controlled ultrastructural features are hydrophilic and mechanically comparable to the sponge quality (6.4 MPa for the compression strength of the PPF graft material vs. 5. for the compression strength of the sponge quality). 0 MPa (Carter and Hayes, Science 194, 1741-1175 (1976))), dimensionally stable, and porous (Bondre et al., 1999). Histological and histomorphological examination of the rat implant area shows that scaffold porosity supports bone ingrowth and that the stability of the scaffold maintains the dimensional integrity of the defect site (Lewandrowski et al., 1999). Preliminary mechanical tests on nano-HA loaded with PPF filler showed that the PPF composite loaded with nano-HA has a compressive strength of 7.5 MPa.

Example 3. PLA-based scaffold with micro-HA or nano-HA for spinal segment repair
For spinal fusion, poly (lactic acid), PLA, fixtures were loaded with HA (either micron or nanosize) and compared to PLA-only fixtures. FIG. 1 shows the effect of the HA filler on the compression properties of this fixture in the spinal segment after L4 / L5 discectomy and anterior and posterior incisions of the longitudinal ligament. FIG. 1 summarizes the missing load and stiffness results normalized with respect to the intact motion segment during axial compression. The polymer only figure (“Bio cate 1”) is 80-100% lower in hardness and defect loading compared to the intact spinal segment, respectively, while “Bio cage 2” filled with micron HA is And 45 to 50% of the value. “Bio cage 3” loaded with nano-HA was statistically equivalent to the unfilled polymer (“Bio cage 1”).

Example 4. PPF-HA for use as a bone graft extender
A PPF-based scaffold formulation was used as a bone graft extender in a rat tibia model. This scaffold was mixed with autografts and allografts and placed directly into a circular metaphyseal defect made in front of the rat tibia using a dental cutter (reaching a diameter of 4.5 mm). This material was cured in situ. These formulations were compared to defects without any of the graft material, autograft, allograft and PPF alone.

(Materials and methods)
Six groups of animals were included in this study. Fresh corticocellous autografts were obtained during surgery and then mixed immediately with PPF bone graft extender (Table 3) prior to reimplantation into the defect site. Allografts of similar nature were obtained from Sprague-Dawley rats, soft tissue and bone marrow were dropped, frozen fresh and stored at −80 ° C. until use. Both autografts and allografts were mixed with a PPF bone graft extender in a 50:50 ratio.

雄性Ｓｐｒａｇｕｅ−Ｄａｗｌｅｙラット（およそ４００グラム、Ｃｈａｒｌｅｓ Ｒｉｖｅｒ Ｂｒｅｅｄｉｎｇ Ｌａｂｏｒａｔｏｒｉｅｓ）を、動物モデルとして使用した。塩酸ケタミン（１００ｍｇ／ｋｇ）およびキシラジン（５ｍｇ／ｋｇ）の筋内注射を使用して、動物に麻酔した。これらのラットに、筋内で予防投与量のペニシリンＧ（２５，０００Ｕ／ｋｇ）も与え、そして外科手術部位を剃毛し、そしてＢｅｔａｄｉｎｅ（ポビドンヨード）およびアルコール（Ｄｕｒａ−Ｐｒｅｐ；３Ｍ Ｈｅａｌｔｈ Ｃａｒｅ、Ｓｔ．Ｐａｕｌ、ＭＤ、ＵＳＡ）の溶液で調製した。１群３頭の動物および６頭の動物を、それぞれ１週目および４週目で屠殺し、それぞれ６群について研究した。従って、合計５４頭の動物を本研究に含んだ。

Male Sprague-Dawley rats (approximately 400 grams, Charles River Breeding Laboratories) were used as animal models. Animals were anesthetized using intramuscular injection of ketamine hydrochloride (100 mg / kg) and xylazine (5 mg / kg). These rats were also given a prophylactic dose of penicillin G (25,000 U / kg) intramuscularly, and the surgical site was shaved and Betadine (Povidone iodine) and alcohol (Dura-Prep; 3M Health Care, St .Paul, MD, USA). Groups of 3 animals and 6 animals were sacrificed at 1 and 4 weeks, respectively, and 6 groups were studied respectively. Therefore, a total of 54 animals were included in this study.

(result)
Histological evaluation of the negative control (not filled with any graft material) showed that the defect was filled with fibrous tissue and granulation tissue 1 week after surgery. At 4 weeks, there was some reactive new bone formation without closing the defect. Defects filled with either PPF alone or autograft bone material alone were used as positive controls. These sections showed early new fibrous bone in the autograft group with nearly complete defect healing at 4 weeks. Defects filled with PPF alone showed early new bone formation, promoted by a PPF cellular scaffold that demonstrates PPF osteoconductivity. An increase in resorbability of this material with continuous, gradual ingrowth of new bone was seen 4 weeks after surgery.

  Mixing the PPF bone graft bulking agent with allograft material or autograft material resulted in enhanced new bone formation with both allograft and autograft. However, improved bone induction was only seen when the PPF bone graft bulking agent was mixed with fresh autograft. The combination of PPF and autograft resulted in more new bone formation than when the autograft was used alone. In addition, there was additional new bone when using allograft combinations versus allografts alone. These histological observations were supported by histomorphometric measurements of new bone formation, which demonstrated a significant increase when PPF was used in combination with either autografts or allografts (Table 4 and 5). Metaphyseal and cortical reconstruction index was determined as an approximate average percentage ratio based on 3 or 6 animals per study group (for the 1 or 4 week time points, respectively). This analysis showed that there was significant additional new bone formation in the experimental group (bone graft mixed with PPF graft bulking agent) when compared to the positive control group (bone graft or PPF graft bulking agent alone). Showed that.

（実施例５．即時性インプラント支持体としての、歯周部組織再生のためのＰＰＦ−ＨＡ泡沫の使用）
（材料および方法）
骨移植片増量剤キャリアを、気泡性パテ（重合可能な実体（ＰＰＦ、ＶＰ）をイニシエーター（ＢＰ）から分離するために、２部系として調製される）として調製した。ＢＰは、重量的には少ない成分であるので、バルクを提供するために過酸化ベンゾイルに関して不活性な成分およびそのインヒビターとともにパッケージする。

Example 5. Use of PPF-HA foam for periodontal tissue regeneration as an immediate implant support.
(Materials and methods)
A bone graft extender carrier was prepared as a cellular putty (prepared as a two-part system to separate the polymerizable entity (PPF, VP) from the initiator (BP)). Since BP is a minor component by weight, it is packaged with a component that is inert with respect to benzoyl peroxide and its inhibitors to provide a bulk.

This foaming agent consists of granules composed of stoichiometric citric acid (CA) and sodium bicarbonate (ie 1.0: 1.3, w / w). The in vitro test specimen was a cylinder prepared by placing the polymerizing composite material in a 10 mm diameter, 10 cm high cylindrical Teflon ^TM mold and cross-linking the polymer foam. . Samples used for this mechanical test were 0.0 mm × 10 mm for in vitro and in vivo tests. The formulation of this example produced a foam having a density of 0.6484 ± 0.0834 g / ml.

Mechanical testing was performed according to ASTM F451-95 for bone cement. This sample was tested during compression on an Instron Model 8511 Materials Testing Machine. The Instron was fitted with a 500 lb load cell and the load weigher was zeroed, balanced and calibrated. The sample was compressed to break at a crosshead speed of 1 cm / min and a load deformation curve was recorded. From these data, the final compressive stress (σ) and Young coefficient (γ) could be calculated. The final compressive stress was calculated as the load applied at failure divided by the original cross-sectional area of the remaining specimen, while the Young's coefficient was calculated as the slope in the linear portion of the load deformation curve. The average compressive strength (average of 8 samples) of the porous cement was 6.77 ± 2.5 MPa. These compressive strength data were comparable to the values reported by Carter and Hayes, who measured this property of bone as a function of strain rate and density. It was noted that at a comparable strain rate, the sponge compressive strength (ρ = 0.31 g / cm ³ ) was 5.0 MPa (Carter and Hayes, 1976).

  Inserting this implant system requires only a brief surgical session that requires a buccal tongue incision made in the gingival tissue in the middle of a toothless period. A procedure for creating a tunnel was then used to separate the tissue from the gingival cleft in both directions from the incision. This was done without further incision of the tissue and without reflection on the tissue. After the tissue separated from the bone, the polymer system was completely injected under the tissue distal to the incision. The inflation system was then shaped in the anterior direction of the incision and placed under the tissue until the desired shape of the periodontal tissue to be augmented or “assembled” was achieved. This molding under the gingival tissue hardened the biodegradable polymer system, resulting in a relatively hard structure.

  Once the periodontal regeneration system is formed over and closely aligned with the gingival cleft of the bone, the incision is sutured so that the operation of the transplant regeneration system is completely completed. During this procedure, a subperiosteal implant system was placed. The passage of time then made it possible to reconstruct this gingival tissue. A post and artificial tooth structure were then placed on top of the implant system placed under the periosteum. To both sides of the incision, the implant system was tunneled, one-half slid under one tissue and the other half slid under one other tissue. The two injected implant systems were then combined and molded into one part.

  Once enough time had passed for the implant part to be firmly fixed in place, a copy of the artificial tooth structure was made. This copy was screwed into the outer thread provided by the implant placed under the periosteum and acted as a fixed posterior crown of the prosthesis.

(result)
Dimensionally stable porous scaffolds were prepared by crosslinking of unsaturated PPF polymer with vinylpyrrolidone (VP) in the presence of sodium bicarbonate and citric acid and hydroxyapatite (HA). To prove feasibility, a PPF scaffold was established as described above, in vitro for morphological, mechanical and surface properties, and a rat tibial defect model established by Gerhart et al. (1989). Were evaluated in vivo. The scaffold has been shown to be mechanically comparable to cancellous, dimensionally stable and porous. Histological and histomorphological experiments of the rat implant area show that this biodegradable bone graft extender scaffold supports the stability of the scaffold maintaining bone ingrowth and the dimensional integrity of the defect site Suggested to do.

  The time sequence of bone ingrowth into the PPF cellular scaffold injected into the tibial drill hole was studied. Osteoclastic and osteoblastic activity and neovascularization were seen at the site of foam implantation as early as 1 week after surgery. The foam appeared to act as a scaffold for new bone formation. At 3 weeks post-surgery, the drilled hole was completely healed in all animals injected with this foam. When compared, this is not the case for control animals where the hole is drilled but the implant is not injected. There was some periosteal bone formation at 3 and 4 weeks after surgery, but complete healing of this hole was not observed in any of the sham-operated animals.

Example 6. Determination of osteoconductivity of porous poly (propyleneglycol-glycol-co-fumaric acid) scaffold
(Materials and methods)
(material)
Poly (propylene fumarate ester) (PPF) is prepared in the presence of p-toluenesulfonic acid (Aldrich) in the presence of fumaric acid (Fisher Scientific, Inc., Pittsburgh, PA, USA) and propylene glycol (Aldrich Chemical Co., Milwaukee). , WI, USA) (Gresser et al., J. Biomed. Mat. Res., 29, 1241-1247 (1995); Lewandrowski et al., Tissue Engineering Tissue Eng, 5 (4): 305-16. (1999)). 1-Vinyl-2-pyrrolidinone (VP), benzoyl peroxide (BP) and NN-dimethyl-p-toluidine (DMPT) were purchased from Aldrich and used as received.

(Bone repair material formulation)
A PPF-based bone graft replacement system was prepared as a two-part formulation consisting of solid powder and liquid components as shown in Table 6. To form a viscous putty-like paste, an aqueous solution of VP (72.6% w / w) and DMPT (0.2% w / w) was combined with PPF (71.8% w / w) and hydroxy. The bone repair system was prepared by mixing into a dry powder mixture of apatite. The weight ratio of PPF: VP was kept constant at 4: 1. The cross-linking reaction between PPF and VP was initiated by the addition of benzoyl peroxide (BP; 3.6% w / w). Free radical production was accelerated by the use of DMPT in the liquid mixture. Sodium bicarbonate (1.7% w / w) and citric acid (1.3% w / w) were also added to the dry powder formulation. Upon mixing of the VP solution and the PPF powder, the reaction of effervescent agents (citric acid (CA) and sodium bicarbonate (SB)) results in controlled swelling of the graft material having a respective pore size of 100-1000 μm. It was.

  Two types of hydroxyapatite (HA) were used to make two specific PPF formulations: sintered, spherical μm sized HA (median particle size = 26 μm, from CAM Implants, The Netherlands). Commercially available) and nm-sized HA (median particle size = 40 nm (Sun T and Ying JY, Nature, 1997, 389: 704-706 (1997)); Ying et al., J. Am. Ceram. Soc. 76 (10) Ying et al. Angew. Chem. Int. Ed. 38 (1): 56-77 (1999); Zhang et al., Chem. Mater. 11 (7): 1659-1665 (1999)). As outlined below, this μm sized HA PP A defect (Group C) filled with a mineral-shedded bone matrix (Group C) and an empty defect that remains unhealed (Group D) with a formulation (Group A) and nm-sized HA PPF formulation (Group B) Compared.

（動物研究の設計）
ＰＰＦベースの骨修復材料の骨伝導効果および生体適合性を評価するために、Ｐｅｔｔｉｓら、Ｊ．Ｏｒａｌ Ｍａｘｉｌｌｏｆａｃ Ｓｕｒｇ ４８（１０）：１０６８〜７４（１９９０）およびＳａｌａｔａら、Ｉｎｔ．Ｊ．Ｏｒａｌ Ｍａｘｉｌｌｏｆａｃ Ｉｍｐｌａｎｔｓ １３：４４〜５１（１９９８）により、以前に記載されたラット頭蓋骨欠損モデルを使用して、処方物を重大ではない頭蓋欠損に移植した。

(Animal research design)
To assess the osteoconductive effect and biocompatibility of PPF-based bone repair materials, Pettis et al. Oral Maxilofac Surg 48 (10): 1068-74 (1990) and Salata et al., Int. J. et al. Oral Maxilofac Implants 13: 44-51 (1998) used the rat skull defect model previously described to implant the formulation into a non-critical skull defect.

The size of this defect reached a diameter of 4 mm. US Public Health Service guidelines for the care and use of laboratory animals were followed. Sprague Dawley rats weighing approximately 100 g and 28 days old were used as animal models (Charles River Laboratories, Wilmington, MA, USA). Animals were anesthetized using intramuscular injection of ketamine hydrochloride (100 mg / kg) and xylazine (5 mg / kg). The surgical site was shaved and prepared with a solution of Betadine (Povidone iodine) and alcohol (Dura-Prep; 3M Health Care, St. Paul, MN, USA). After surgery, these rats were also given a prophylactic dose of penicillin G (25,000 U / kg) intramuscularly.
Two 4 mm diameter cortical defects were made in the skull of each rat. The animals were grouped into 3 groups with 8 animals per group, and one deficiency was treated with one of the following: PPF formulations containing μm size HA (Group A), nm size HA and PPF ( Group B), or bone matrix with mineral loss (Group C). This second defect was left unhealed and served as a mating control. The group of animals was evaluated at 1 week, 2 weeks, 4 weeks, and 7 weeks after surgery, so a total of 96 animals were included in the study. The PPF formulation was mixed during surgery to be similar in hardness to a paste or putty and then implanted into the prepared skull defect using a spatula. The bone repair material formulation was cured in situ. For implantation in an animal group C, the bone matrix and demineralized ^{(DMB, Grafton Putty TM, Muskuloskeletal} Transplant Foundation, Shrewsbury, NJ) was obtained. The implant material was allowed to cure in situ for approximately 5 minutes, and then the soft tissue and skin were closed in layers with resorbable sutures.
(Evaluation method of bone repair material formulation)
After sacrifice, excisional biopsies of the skull and surrounding soft tissue were taken on radiographs. The specimens were then fixed in 10% neutral buffered formalin and decalcified in 4N formic acid. Serial longitudinal sections 5 μm thick were made at 50 μm intervals. Sections were stained with hematoxyli-eosin (H & E). Slides were tested for defect healing by descriptive analysis of resorption activity and new bone formation in cranial defect bones, and for inflammatory responses to bone repair materials. Furthermore, by using a CCD video camera system mounted on a Zeiss microscope (TM-745; PULNiX, Synnyvale, CA, USA), images of serial longitudinal sections of the specimen were obtained and Histomorphometry evaluation of new bone formation according to transplantation of PPF repair material was performed. Images were digitized and analyzed using Image Pro Plus software. The area occupied by new bone within this defect was quantified using H & E stained slides from 2 animals at 1 week, 2 weeks, 4 weeks and 7 weeks. This new bone formation (expressed as a percentage of the original 4 mm defect area compared to each animal's empty control) was converted into three templates, or areas of the target mask (this skull defect and Was measured for each sample using 3) located in three regions over the corresponding opposite region. Averages were obtained for each sample from a minimum of 3 consecutive longitudinal sections and a maximum of 6 consecutive longitudinal sections. This makes it possible to obtain an approximate absolute volume of newly formed bone (defined as New Bone Volume Index), which is the average of these continuous area measurements for each bone specimen. It is given as (mean value ± standard deviation). The New Bone Volume Index is given as a percentage ratio and is shown as the average of all sections prepared from animals transplanted with PPF per group.
(Statistical analysis)
Differences in the amount of new bone formed in response to transplantation of PPF formulations containing either μm size HA (Group A) or nm size HA (Group B), ANOVA test for normally distributed samples Was used to analyze for statistical significance. A Tukey test was used to establish the P value. Samples that were not normally distributed were analyzed using the Kruskal-Wallis test. A p level of less than 0.05 was considered statistically significant.

(result)
There were no postoperative complications or clinical signs of implant response at 96 experimental implantation sites. No fractures or deep infections were observed throughout the postoperative period. After dissection and incision and macroscopic examination before embedding for histological and morphometric analysis.

  The results are described in Table 7 and Table 8.

全ての移植した標本は、種々の量の新たに形成された骨で充填されていることが見出された。空の欠損部位は、見出されなかった。全てのＰＰＦおよびＤＭＢグラウトされた骨標本を、インタクトなまま回収した。ＰＰＦベースの骨修復材料材料の頭蓋骨欠損部位への移植は、回収した標本のいずれにおいても、過剰な肉眼で見える肉芽組織形成の非存在によって証明されるような、全体にわたる良性組織応答を生じた。全ての手術部位は、十分に治癒したようであり、そしてインサイチュで硬化された材料に対する周囲の柔組織の明白な有害反応は存在しなかった。コントロール欠損の全ては、良性らしい柔組織反応を有した。

All transplanted specimens were found to be filled with various amounts of newly formed bone. An empty defect site was not found. All PPF and DMB grouted bone specimens were collected intact. Implantation of PPF-based bone repair material into the skull defect site produced an overall benign tissue response, as evidenced by the absence of excessive gross granulation tissue formation in any of the recovered specimens. . All surgical sites appeared to heal well and there was no obvious adverse reaction of the surrounding soft tissue to the in-situ hardened material. All of the control deficits had benign soft tissue responses.

(X-ray photography research)
X-ray analysis of all experimental skull specimens at different follow-up time points showed sufficient evidence of bone treatment at the implantation site up to 4 weeks after surgery, regardless of which implant material was used. It was shown to exist. There was no radiographic lucities in and around the defect site. At 4 weeks, radiographs showed more bone formation in the two PPF-based groups than in the DMB group. There was no evidence of bone formation in the control defect left unhealed. The amount of new bone formation was evaluated semi-quantitatively using light absorption measurements, and this amount was normalized between radiographs with internal phantoms. This measurement showed the greatest amount of new bone formation in a PPF implant containing nm sized HA. On radiographs, the surrounding parenchyma was normal in appearance and there was no evidence of swelling or fluid retention at all implantation sites.

(Histological analysis)
In Group A, a PPF-based bone repair material containing μm size HA was implanted. Histological analysis of bone samples collected at 1 and 2 weeks post-operatively showed that most of the bone repair material hardened in situ remained intact. There were several new bone formations, which occurred centrally from the surface of this defect in an afferent manner. At 4 and 7 weeks post-surgery, the PPF implant appeared to be increasingly replaced by newly formed bone. The PPF implant was no longer intact and it seemed that degradation following bone ingrowth occurred. It was noted that in most samples the defect was healed but not completely filled with new bone.

  In group B, a PPF bone repair material containing nm size HA was implanted. Histological findings at 1 and 2 weeks after surgery were similar to the group A samples. However, bone formation appears to be more active with a more pronounced early inflammatory response, as evidenced by bone marrow proliferation achieved within the newly formed bone palisade around the implant. At 4 weeks post-surgery, the nm-HA / PPF implant was completely resorbed and replaced with newly formed bone. At 7 weeks, the cranial defect was essentially healed using near-perfect reconstruction of the defect area when compared to the sample obtained 4 weeks after surgery.

  In group C, the mineral-depleted bone matrix was transplanted for control purposes. Histological observations differed from the groups transplanted with PPF (Groups A and B) in that the transplanted and mineral-depleted bone matrix could not be clearly identified at any postoperative follow-up time point. It was. This material appeared to be resorbed as early as the first week after surgery. However, new bone formation was noted as early as 4 weeks after surgery and appeared to be more apparent at 7 weeks after surgery. Near complete cranial defect healing was noted 7 weeks after surgery. Seven weeks after surgery, the entire cranial defect was filled with newly formed bone that was loosely packed.

  In contrast, the control defect (where no implant was placed) remained empty until 4 weeks after surgery. Several reactive bone formations were noted, originating from the surface of the drilled defect. At 7 weeks post-surgery, the control drilled defect appeared similar to the 4 week sample and was not filled with newly formed bone.

  This was supported by histomorphometric analysis. By quantitative volume measurement expressed by New Bone Volume Index, experimental defects treated with nm-HA / PPF implants showed a maximum amount of new bone formation (mean value 95 ±) compared to control defects (no implants). 17 percent) (p <0.02).

  In the reconstruction of the maxillofacial and mandible, filling the bone void with the graft material remains difficult due to lack of structural support during the process of new bone regeneration. Of unsaturated polyester poly (propylene glycol-co-fumaric acid) (PPF) and two different types of hydroxyapatite (either μm-sized or nm-sized hydroxyapatite filler, and foaming foam A bioresorbable bone repair material (crosslinked in the presence) was prepared. This bone repair material develops porosity in vivo by generating oxygen dioxide during the reaction of citric acid and sodium bicarbonate, the reactions of these materials each have a pore size of 100-1000 μm, Responsible for controlled pore generation and expansion. Two, unimportant, 4 mm diameter cortical defects were made in the calvaria of 28 day old Sprague Dawley rats. In each animal, one defect was treated with one of the following materials: a PPF formulation with μm sized hydroxyapatite, a PPF with nm sized hydroxyapatite, and a mineral shed bone matrix. To serve as a pairing control, the second defect was left unhealed. Four sets, each containing 24 animals, were evaluated at 1 week, 2 weeks, 4 weeks, and 7 weeks after surgery, with 8 animals treated with each filler material at each evaluation. used. Radiographic and histological techniques were used to analyze the amount of new bone formation and the presence of inflammatory infiltrates at the mating site.

  Histological analysis of the healing process showed superior healing of the skull defect using both formulations of PPF bone repair material when compared to the control defect (which was left empty). The remodeling of newly formed bone appeared to be more advanced using a PPF formulation containing nm-sized hydroxyapatite. These findings were confirmed by histomorphometric analysis of new bone formation. Inflammatory cells were noted only in the week 1 group and were not related to the type of material. Inflammatory infiltration was absent in all other groups assessed at the latter postoperative time. The results of this study demonstrated both the biocompatible and osteoconductive properties of this porous PPF-based bone repair material in a skull defect model.

Example 7. Double-blind and controlled parallel study of four PPF / HA formulations for healing of mandibular defects in rats
The subject of this study was to assess the efficiency of dysplastic graft formulations administered topically to promote healing of mandibular defects created in rats.

(Materials and methods)
160 rats were randomized into 5 groups with 32 animals per group: positive control group and 4 test groups. All groups were assigned different treatments. On day 0, the health status of each animal was examined. These animals were then weighed, randomized and numbered. Surgery was performed. Specimens were collected at 4 time points (week 1, week 2, week 4 and week 7).

  Animals were weighed and then anesthetized with a single intraperitoneal injection of ketamine hydrochloride (100 mg / kg) and xylazine (5 mg / kg). Mandibular branches were exposed symmetrically with respect to the right and left sides. In each animal, a mandibular defect was created on each side by drilling a 4 mm hole using a rotary drill. The defect was washed with Ringer lactate. In all animals, this left defect was not treated. The defect on the right side of group 1 was packed with mineral-shedded human bone matrix, commercially available as Grafton Putty® from Musculoskeleton Transplant Foundation. The defects on the right side of Groups 2, 3, and 4 were packed with one of three PPF-based formulations, and Group 5 was packed with autograft only. The incision was closed using a knot suture. This area was cleaned and covered with a wound healing liquid. The rats were given a peritoneal prophylactic dose of penicillin G (25,000 U / kg). One hour after surgery, each rat received an intraperitoneal dose of buprenorphine (0.05 mg / kg).

  Eight animals from each group were sacrificed at 1 week, 2 weeks, 4 weeks, and 7 weeks after surgery. The specimens were wrapped in gauze soaked in saline and sent to the sponsor for evaluation.

  Test materials (crosslinked PPF containing HA particles, organized as described above) were obtained from Cambridge Scientific, Inc. Prepared by a technician from The test material was packed slowly (mixed within 2 minutes) with low viscosity into a defect made in the right mandible. In a group of positive controls, a combination of powder and saline was used to mix mineral-depleted human bone matrix (Grafton Putty® from Musculoskeleton Transplant Foundation) into a high viscosity slurry and then a slight overfill. Packed in a defect with The defect was closed 5 minutes after placement of the test implant material. The incision was closed with a nodule suture.

  Sectioned mandibles (2 per film) were placed on the outside on Kodak Occusal film. Radiographs were taken. Radiographs of the left and right defects were blinded for treatment and measured along the x and y axes. The average of the x-axis and y-axis measurements was calculated and the total area was calculated. The total area was graphed. The percent difference between the left defect and the right defect was calculated and graphed.

(result)
The average mandibular defect area for each test material was determined. The average area of the left defect (untreated control) and right defect (test implant) was determined and this was significant difference between 100% autograft and its control at week 4 (p < 0.001) and a significant difference (p = 0.007) between the 75% PPF / HA + 25% autograft and its control at 4 weeks.

Area plots for untreated defects and defects that received mineral deciduous bone from the same mandible were calculated from the mean values of the x and y axes using the area of the circle equation (πr ² ). The group mean and the standard deviation (SEM) of this mean were calculated for each week. This plot shows the area for untreated defect areas and defects that receive mineral-depleted bone.

Mandibular defect area treated with PPF / HA formulation was determined. The area for the untreated defect and the defect receiving 100% PPF / HA from the same mandible was calculated from the average of the x and y axes using the area of the circle equation (πr ² ). Group mean and standard deviation (SEM) of this mean were calculated for each week for the untreated defect and the defect receiving 100% PPF / HA. FIG. 5 compares the defect area for untreated defects and defects receiving 25% PPF / HA-75% autografts and defect areas for untreated defects and defects receiving 100% autografts. The t-test shows that the untreated defect area is significantly smaller (p <0.001) than the defect area transplanted with 100% autograft at 4 weeks. Areas for untreated defects and defects that received 75% PPF / HA-25% autografts were compared. The t-test showed that the untreated defect area was significantly smaller (p = 0.007) than the area of the defect transplanted with 100% autograft at 4 weeks. These results demonstrate that the untreated (negative) controls of the four test groups are statistically equal, except for the mineral depleted bone and 100% PPF / HA at week 1. These results show that the PPF / HA mixture achieves healing at a rate similar to treatment with 100% autograft with or without autograft.

(Example 8. Repair of skull defect)
A similar study was carried out to repair the skull defect.

(Method)
In each animal, two skull defects were created using an established animal model. Prior to surgery, animals were anesthetized by intraperitoneal injection of ketamine hydrochloride (100 mg / kg) and xylazine (5 mg / kg). The surgical site was shaved and polished with a solution of Betadine (Povidone iodine) and alcohol (Dura-Prep; 3M Health Care, St. Paul, MD, USA). The calvaria was exposed by making a 3 cm longitudinal incision in the rat cranium. The periosteum was peeled off and two 4 mm diameter full-thickness bone holes were made in parallel by removal of similarly sized bone discs using a rotary drill and washed with lactated Ringer's solution. Transplantation began after hemostasis was achieved. In all animals, the left defect was untreated. The right defect received the implant. The soft tissue and skin were closed in layers with intermittent absorbable sutures. These rats were given a peritoneal prophylactic dose of penicillin G (25,000 U / kg). One hour after the operation, buprenorphine (0.05 mg / kg) was intramuscularly administered as an analgesic.

(result)
The following results were obtained:
The defect area filled with mineral-depleted bone is 8.2% smaller than the control defect area in the first week. The defect area filled with mineral-depleted bone is 0.7% larger than the control defect area at 2 weeks. The defect area filled with mineral-depleted bone is 56.1% smaller at 4 weeks than the control defect area. The defect area filled with mineral-depleted bone is 117.8% smaller than the control defect area at week 7.

  The defect area filled with PPF / μm HA is 30.7% smaller at week 1 than its control defect. The defect area filled with PPF / μm HA is 29.4% smaller than the control defect at 2 weeks. The defect area filled with PPF / μm HA is 150.6% smaller than the control defect at 4 weeks. The defect area filled with PPF / μm HA is 307.4% larger than the control defect at 7 weeks.

  The defect area filled with PPF / nm HA is 23.6% smaller at week 1 than the defect in the control. The defect area filled with PPF / nm HA is 30.5% smaller at week 2 than the defect in the control. The defect area filled with PPF / nm HA is 123.0% smaller than the control defect at 4 weeks. The defect area filled with PPF / nm HA is 30.4% smaller than the control defect at 7 weeks. At week 2, there was a significant difference in defect area in defects filled with PPF / nm HA versus untreated defects (p = 0.004). At week 2, there was a significant difference in the defect area between defects filled with PPF / nm HA versus defects filled with mineral-shedded bone (p = 0.014).

FIG. 1 summarizes incomplete loading and stiffness relative to an intact motion segment during axial compression. FIG. CSI-02. Mandibular defect area. The area of the untreated defect and the defect that received bone that had been mineralized from the same mandible were calculated from the mean values of the x-axis and y-axis using the area of the circular equation (πr ² ). Group means and standard deviations (SEM) were calculated each week. This plot shows the area for defects that receive untreated and mineral deciduous bone.

Claims (32)

A bioresorbable osteoconductive composition for bone repair, comprising:
a) a bioresorbable polymer b) a biocompatible micro- or nano-filler; and c) a pore or pore-generating material,
A composition comprising: 2. The composition of claim 1, wherein the filler is selected from the group consisting of metals, calcium carbonate, carbon, bioceramics, and synthetic materials. 2. The composition of claim 1, wherein the filler is a bioceramic. The composition of claim 1, wherein the filler is hydroxyapatite. The composition of claim 1, wherein the filler is a metal. 6. The composition according to claim 5, wherein the bioresorbable polymer is poly (L-lactic acid), poly (DL-lactic acid), poly (DL-lactic acid-co-glycolic acid). ), Poly (glycolic acid), poly (ε-caprolactone), polyorthoester, polyanhydride, polydioxanone, copoly (ether-ester), polyamide polylactone, and citric acid, isocitric acid, cis-aconitic acid, α -A composition selected from the group consisting of polyesters produced from acids selected from the group consisting of ketoglutaric acid, succinic acid, malic acid, oxaloacetic acid and fumaric acid, and combinations thereof. 7. The composition of claim 6, wherein the bioresorbable polymer is a polyester that is crosslinkable using a crosslinker. 8. The composition of claim 7, wherein the bioresorbable polymer is poly (propylene glycol-fumaric acid). 9. A composition according to claim 8, wherein the cross-linking agent is selected from the group consisting of vinyl pyrrolidone and methyl methacrylate. 10. The composition of claim 9, wherein the cross-linking agent is vinyl pyrrolidone. 2. The composition of claim 1, comprising pores in the range between 100 microns and 1000 microns. The composition of claim 1 comprising a pore former, wherein the pore or pore-generating material is a foaming agent. 13. A composition according to claim 12, wherein the foaming agent is a carbonate and an acid. 2. The composition of claim 1 comprising a pore-forming agent, wherein the pore-forming agent is selected from the group consisting of particles and ammonium carbonate that are leached for the polymer using a non-solvent. . The composition of claim 1, comprising poly (propylene glycol fumaric acid) and hyaluronic acid particles as fillers. The composition of claim 1, further comprising a biologically active material. The composition of claim 1, further comprising a graft material selected from the group consisting of allografts, autografts, xenografts and mineral deciduous bone. 17. The composition of claim 16, wherein the biologically active material is selected from the group consisting of cells and therapeutic agents. 19. The composition of claim 18, wherein the biologically active material is selected from the group consisting of growth factors, antibiotic materials, antiviral agents, antifungal agents, immunostimulatory agents, and immunosuppressive agents. A composition which is a therapeutic agent. 2. The composition of claim 1, wherein the composition is in the form of a scaffold. 2. The composition of claim 1, wherein the composition further comprises a fiber or other structural support. 2. A composition according to claim 1, wherein the composition is formed into an implant for implantation into a site for bone repair or bone regeneration. A method for bone repair or regeneration, comprising:
Providing the composition to the site where they are needed,
Wherein the composition is
a) a bioresorbable polymer b) a biocompatible micro- or nano-filler; and c) a pore or a pore-generating material as defined by any one of claims 1-22.
Including a method. 24. The method of claim 23, wherein the composition is implanted at a site for periodontal repair or regeneration. 19. The method of claim 18, wherein the composition is implanted at a site for alveolar repair or regeneration. 19. The method of claim 18, wherein the composition is implanted at a site for repair or regeneration of the maxilla. 21. The method of claim 19, wherein the composition is implanted at a site for periodontal regeneration, wherein the implanting step comprises:
Inserting a periodontal implant onto a ridge of a mammalian subject, further comprising:
a) making an incision in the tissue covering the bone ridge;
b) tunneling the tissue such that the tissue separates from the bone ridge;
c) injecting a periodontal regeneration system under the tissue;
d) fixing the two parts of the implant together;
e) suturing the incision; and f) inserting a subperiosteal implant system;
Including the method. 28. The method of claim 27, comprising administering a bioactive agent with the composition, and applying a therapeutically effective amount of a growth factor directly to the periodontal regeneration system. 30. The method of claim 28, wherein the growth factor is a platelet-derived growth factor in the form having two beta chains (PDGF-BB), a platelet-derived growth factor in the form having an alpha chain and a beta chain. (PDGF-AB), IGF-I, and a method that is one or more factors selected from the group consisting of TGF-β. 24. The method of claim 23, wherein the composition is formed into a prosthetic implant prior to implantation. 24. The method of claim 23, wherein the composition forms a bone graft extender. 32. The method of claim 31, wherein the composition comprises a graft material.

Images